gms | German Medical Science

A variety of material properties like hardness, shape, edge finish, porosity and site of implantation influence the biological response after placing an implant in the body.

Surface properties are very important in the areas of adhesion and bondability and influence biological responses. Contact angle measurement techniques have evolved from theoretical and empirical approaches to a rapid means to gain information about the surface constitution of materials and its role in interfacial phenomena.

One advantage of contact angle methods in the study of surface interactions is that they probe only the outermost layer (10–100 Å, 1–10 nm) of a surface and are capable of testing hydrated specimens. A limitation of contact angle testing is that the test environment is still artificial; thus, results may not correlate well with clinical observations. A liquid placed on a solid will not wet the solid but remains a cohesive drop, having a definite angle of contact between the liquid and solid phases. This contact angle can be related to surface constitution through the examination of the physical and chemical forces involved, first described qualitatively by Young in 1805. This basic relationship that the spreading of a drop on a surface is related to the physical-chemical forces of the liquid, the solid and the environment forms the basis of all contact angle measurements.

The application of Young’s equation is extended to define various physical-chemical parameters like predicting material wettability and different test systems. Two primary attraction forces cause molecules to associate into liquids and solids: the dipple-dipole forces of polar molecules (also including hydrogen bonding) and the London (dispersion) forces that affect all molecules. The interfacial free energy of a polymer in its environment is the result of polar and dispersive forces, and it represents unfulfilled bonding of the surface in that environment.

We present detailed contact angle data on a variety of different intraocular lens (IOL) materials including polymethyl methacrylate (PMMA), heparin-surfacemodified PMMA, silicones and foldable acrylics, both hydrophilic and hydrophobic. By making this information generally available, we hope to advance the understanding of the relationship between IOL materials, surface properties and biological interactions in vivo.

The most common contact angle method is measuring the contact angle of a water droplet on a surface in air, the sessile drop method. Our experiments showed that surface properties and information obtained by contact angle methods can vary widely depending on the material used. To avoid ambiguity, the contact angle measurement method should always be specified, because the captive bubble measurement method is used in very few studies.

Surface roughness can influence contact angle measurements. In this study, we found the optic surfaces where the contact angle measurements were made to be smooth at a magnification of 3000. Thus, in this series the roughness of the surfaces was comparable.

Surface cleanliness can also influence contact angle measurements because contaminants can alter the readings. All IOLs were very clean. Thus, surface contamination was minimal and unlikely to significantly influence the contact angle results.

Studies have been done to determine the contact angle dependence on surface curvature. In general, curvature effects can be neglected when the drop’s radius is larger than its penetration thickness. The penetration thickness reflects the amount that the droplet moves into the material. This can be conceptually appreciated for hydrogels or hydrophilic coatings, where a water droplet would eventually be absorbed into the material. The water droplets that were used were not absorbed into any of the hydrophobic materials to any extent. The reading was, however, taken seconds after the droplet had been placed on the surface. Moreover, all IOLs had a close IOL power range. Therefore, because of the types of materials used in this study, the curvature effects were negligible.

The common classifications of hydrophobic and hydrophilic arose from the sessile drop method. In our study, all materials tested appeared to be more or less hydrophobic, as demonstrated by their high water contact angle reading. The PMMA (AMO DuraLens PS 101A), AcrySof (Alcon), the Siflex 4 and the Flex 60 have similar water contact angles (73.2, 73.3, 75.4 and 75.7°, respectively). The AMO soft acrylic Sensar had a somewhat higher water contact angle (81.7°). Using the sessile drop method, we saw the influence of the trifluoroethylene component of the material. In the hydrophobic environment of the air, the fluorinated component of the Sensar material surfaced and thereby increased the water contact angle of the material so that its value was somewhat higher than that of PMMA. We saw relatively small differences between most materials with highest values for the heparin-surface-modified material; that is, critical surface tension values for these materials were similar. Thus, we would expect these materials to exhibit similar wetting behaviors for a given liquid. The surface tension of water is 72.1 dyn/cm. From these data, we can predict that the wetting behavior of all these materials is similar, that is none of these materials is really wettable by water. For these surfaces to be wettable, the critical surface tension of the material must be greater than the surface tension of the liquid. The chain mobility of the polymer surface for silicones and soft acrylics is evident by comparing the water contact- angle-in-air reading with the air bubble contactangle- in-water reading. This mobility has been recognized in other biomaterial applications and is believed to be caused by primary (segmental backbone) and secondary (side chain) molecular motions.

In a previous study, the greatest molecular flexibility was observed in the polydimethyl siloxane, in which the difference between the air and water contact angle readings was 46.7°. The other silicones and flexible acrylics exhibited differences of 23–29° under these conditions. It is important to consider the effects of the environment and molecular mobility when correlating surface properties to biological interactions. Siloxane materials, although they have a polar siliconoxygen bond, have minimal polar interactions. This is because of the symmetrical arrangement of the siliconoxygen bonds and the side groups, and any deviations from perfect symmetry reflect the influence of cross-linkers and reinforcing agents. Although PMMA and acrylics have a large polar component than silicone elastomers, the values are still small. Examination of their chemical structure shows that they have a nonpolar backbone but a polar ester group. This side group polarity is reduced by the substitution on the ester group and the overall symmetrical nature of the polymers. Although their polar components are small, all these materials have the ability to exhibit some level of polar interactions in a polar environment. The polar forces are permanent dipoles, while the dispersion forces are temporary dipoles induced by the environment. Dispersion forces exist in all molecules. These forces arise from temporary dipole moments induced by nearby molecules. If dispersion forces are strong, molecules will be significantly attracted to one another. Dispersive forces are also stronger when intermolecular attractions are maximized through the total contact area between polymer chains. The dispersive component of the surface measures the interactions in a nonpolar environment. This type of environment is not usually physiologically relevant, except in unique circumstances. Silicone oil tamponade procedures are a case in point. In such cases, vitreous replacement in a pseudophakic eye with a posterior capsulotomy may cause the IOL to be exposed to a nonpolar environment. The dispersive components are highest for the polydimethylsiloxane-based lenses (e.g., Bausch & Lomb Soflex 2, Bausch & Lomb Silens 6). By using the contact angle and surface tension to solve the equations derived by Andrade et al., the dispersive components are more intermediate for the Sensar (AR-40e, AMO), AcrySof (Alcon), and several PMMA IOLs.

In a study by Cunanan et al., the dispersive components were lowest for the second-generation siliconesand not very different from the dispersive component of water, 21.6 dyn/cm. The polydimethyl siloxane materials had the highest dispersive forces because they are linear molecules with a high molecular surface area. Because of their linear structure, these molecules overlap well and have many intermolar interactions between the polymer chains. The ability to induce temporary dipole moments can be expected to be greatest with these materials. The polydimethyldiphenylsiloxane material (SI-40NB) had the lowest dispersive component of all materials because of the smaller molecular surface area and the phenyl side groups close to the polymer backbone, which reduce the amount of intermolecular contacts and thus the dispersive forces. The dispersive forces in PMMA and the acrylics were intermediate because of their longer side chains and methyl substitutions off the polymer backbone, which somewhat reduce their intermolecular interactions. The interfacial free energy represents unsatisfied bonding of the surface in the environment tested.

There have been numerous attempts to correlate surface properties with biological interactions, with varying degrees of success. It is well known that upon immersion in biological fluids, all biomaterials become coated with proteins, and extensive studies have been done to document and model these interactions. Cell and tissue adhesion have also been related to surface properties, and attempts have been made to directly measure surface properties of proteins and cells.

These efforts have been largely directed toward understanding the factor(s) responsible for the observed biological interaction. Although a given study may explain the observed biological interactions in that particular study, these factors frequently cannot be generalized to explain other different, but related, interactions. It is likely that biological interactions are influenced by numerous factors, including material surface properties and other parameters. It is becoming increasingly recognized that proteins and cells interact through specific, receptor-mediated interactions. The study of material surface properties alone cannot satisfactorily explain these specific interactions. Studying material surface properties is still important but it must be recognized that additional, specific biological reactions are also involved in observed biological phenomena. In this study, we investigated only material surface properties, but did not evaluate any correlations between biological reactions and material surface properties. On the other side, implantation of IOLs with a low water contact angle and a high critical surface tension like the heparin-surface-modified IOLs were associated with a reduced inflammatory response for routine glaucoma, uveitis and diabetes patients when compared with non-heparin-surface-modified lens implants. Binding heparin to the PMMA surface appears to decrease foreign-body-induced inflammation by making the IOL more hydrophilic.

Bacterial adherence to the IOL may play an important role in the pathogenesis of intraocular infections after cataract surgery. Several studies demonstrated that bacterial adherence is influenced by IOL materials. Bacterial adherence to IOL surface occurs in two phases. The first phase, dependent on the physical characteristics of both micro-organisms and biomaterials and involving reversible attraction caused by electrostatic and Van der Waals forces and hydrophobic bonding, takes place immediately. The second phase is mediated by the bacterial production of a polysaccharide glycocalyx on the IOL surface. Therefore, hydrophobic IOLs, like the AcrySof IOL (water contact angle: 73°) or silicone IOLs, are more permissive to bacteria like Staphylococcus epidermis than hydrophilic IOLs, like the heparin-surface-modified PMMA ones (water contact angle: 57°). Recent studies have shown that silicone IOLs (water contact angle: about 150°) are more permissive to staphylococcal adherence than PMMA IOLs (water contact angle: about 75°). Nevertheless, there is not enough experimental evidence to conclude that the implantation of an AcrySof IOL or silicone IOL is associated with an increased risk of postoperative endophthalmitis. It is important to know that different bacteria show different degrees of adherence to different IOL materials. AcrySof IOLs are less susceptible to primary adherence and biofilm formation by Pseudomonas aeruginosa than PMMA and silicone IOLs. S. epidermidis appeared in preference to be colonizing sites with surface irregularities. Scanning electron microscopy of each lens type investigated in this study revealed consistently smooth surfaces. In vivo, rapid protein absorption may affect the final properties of the IOL. The host reaction could overcome a bacterial infection in its initial phase and prevent the development of an endophthalmitis.

The contact angle technique can yield information that can be used to understand some aspects of material interactions. To date, most relationships are still made empirically by correlating observed biological phenomena to measured surface properties. Biomaterial development has yet to advance to the predictive stage, at which complex interactions with cells, proteins and tissues can be modeled mathematically or with simple in vitro test methods and material properties can be designed to direct specific reactions.